The need and demand for an accurate, non-invasive method for determining blood glucose level in patients is well documented. Barnes et al. (U.S. Pat. No. 5,379,764) disclose the necessity for diabetics to frequently monitor glucose levels in their blood. It can be recognized that the more frequent the analysis, and the more analysis can accurately predict future glucose levels, the more readily large swings in glucose levels can be controlled or eliminated. These large swings are associated with the symptoms and complications of the disease, whose long-term effects can include heart disease, arteriosclerosis, blindness, stroke, hypertension, kidney failure, and premature death.
Presently, most home glucose monitoring is limited to methods and systems requiring a lancet cut into the finger. Repeated lancet cuts lead to repeated irritation and inconvenience, discouraging vigilant self-monitoring. This is believed so compromising to the diabetic patient that the most effective use of any form of diabetic management is rarely achieved.
Infrared spectroscopy is a generally known basis for analysis. In some forms, infrared spectroscopy measures the electromagnetic radiation (typical wavelengths in the range of 0.7-25 μm) a substance absorbs at various wavelengths, though other methods measure other effects a substance has on incident light. Absorption phenomena can include induced molecular vibrations and shifts in energy levels of individual atoms. Either phenomena causes the absorbing molecule or atom to switch to a higher energy state. These phenomena can be statistically shown to occur most frequently in limited ranges of wavelengths. Thus, for light passing through a substance at several wavelengths, the substance will absorb a greater percentage of photons at certain wavelengths than it will at others. The characterization of substances by their spectral absorption characteristics is well known.
At the molecular level, many primary vibrational states transitions occur in the mid-infrared wavelength region (i.e., wavelengths between 3-6 μm). However, non-invasive analyte detection in blood in this region is problematic, if not impossible, because water makes up the majority of the blood and demonstrates strong absorption characteristics at mid-infrared wavelengths. Typically, the problem is overcome through the use of shorter wavelengths that are not as strongly attenuated by water. Frequency overtones of the primary vibrational states exist at higher frequencies (thus shorter wavelengths), enabling quantitative determinations at these wavelengths.
It is known that glucose absorbs at multiple frequencies in both the mid- and near-infrared range. There are, however, other analytes in the blood which also absorb at similar frequencies. Due to the overlapping nature of these absorption bands, reliable non-invasive glucose measurement would be very difficult if only a single frequency were used for analysis. Analysis of spectral data for glucose measurement is facilitated by evaluation of absorption characteristics at several wavelengths, enabling the sensitivity, precision, accuracy, and reliability necessary for quantitative determination. Similar observations apply to other analytes such as urea and ethanol.
An additional difficulty can arise with respect to the incident power required to achieve a readable signal with respect to glucose levels. Glucose is a minor component by weight in blood, making high sensitivity an important aspect of design. One way to achieve higher sensitivity is to use a relatively large input power, which, however, can cause non-linearities in the optical systems themselves as well as in the material subjected to optical radiation. An alternative solution is to use measurements at several wavelengths and incorporate multivariate analysis techniques to statistically “filter out” data not related to the blood glucose concentration. Such techniques can be likewise applied to determine concentrations of other analytes, for example, constituents such as urea.
Many diabetic patients must lance themselves four to five times per day in order to measure their capillary blood glucose concentration and adjust insulin therapy and meals. Optical measurement provides an alternative means for measuring blood glucose by using the absorption or reflection spectra of tissue as a surrogate for the spectra of drawn blood. However, a problem encountered in non-invasive skin based measurements of standard medical blood analytes is the disparity between the concentration of a given analyte in the blood and the same analyte in the adjacent skin tissue water or interstitial fluid.
Blood, which, like tissue, is approximately 80% water, makes up less than 10% of the tissue volume. Since glucose is not made, but only disposed of, in skin, all of the glucose in the water that bathes cells (interstitial fluid) and that is inside cells comes from the blood vessels. That is, blood glucose must move out of the blood vessels and into the surrounding interstitial water and then into cellular elements. This effect is time dependent as well as dependent upon recent and present concentration gradients, the relative juxtaposition of compartments making up the tissue and the adjacent blood vessels, and the relative blood flow to the tissue. Thus, a measurement of total tissue glucose concentration is often very different from the concentration of glucose in the small blood vessels that make up a fraction of the total tissue volume.
Since glucose is only degraded in the skin (not manufactured), the interstitial space must be “filled” with glucose by the local blood vessels. As with any filling process, this is time dependent. Time lags between the concentration of glucose in interstitial fluid and blood have been documented ranging from 0 to 60 minutes, with an average lag of 20 minutes. Thus, the fact that the glucose must move between the tissue and blood causes errors in both interstitial space glucose and total tissue glucose concentration measurements when used as surrogates for blood glucose concentration measurements.
When measurements of total tissue or interstitial glucose concentration and blood glucose concentration are made concurrently, the two are correlated, but the tissue glucose concentrations lag behind the blood levels. Blood or serum glucose concentrations must be delayed in order to overlay the interstitial or total glucose concentration. When blood glucose concentration is changing rapidly (as might be expected in a diabetic after a meal high in sugars or after an insulin injection), the delay is more obvious and the difference between the blood glucose concentration and the other two measurements is most pronounced. Other analytes, such as ethanol, have similar tissue-blood kinetic behavior. The specific magnitude of the “lag” between the tissue and blood analyte concentrations is generally a function of the analyte's specific diffusion properties and each subject's physiology.
This presents obvious problems with respect to using the surrogate methods for monitoring and guiding therapy in diabetic patients. Given the concentration difference, determining whether a given technique is working based on infrequent, discrete measurements is believed impossible. Without more frequent measurements, it is difficult to determine whether the patient's blood glucose is in a steady state condition or is in flux.
The worst-case scenario in diabetic glucose management would be a quickly falling blood glucose concentration. Such a situation can result following a large insulin injection, unopposed by either glucose production in the liver or carbohydrate uptake from food in the digestive system. If a tissue measurement were made it would inappropriately report a level that is higher than the actual blood glucose concentration and that would not inform the patient of the quickly changing concentration. The patient would be unaware of their rapidly falling blood glucose level. The result of very low blood glucose concentrations (below 40 mg/dl, 2.2 mmol) is often loss of consciousness, coma and even brain damage or death if the patient is not discovered in time for medical intervention. For this reason, the ability to detect changing blood glucose levels is greatly desired.
Robinson et al. (U.S. Pat. No. 4,975,581) disclose a method and apparatus for measuring a characteristic of unknown value in a biological sample using infrared spectroscopy in conjunction with a multivariate model that is empirically derived from a set of spectra of biological samples of known characteristic values, and is incorporated herein by reference. The above-mentioned characteristic is generally the concentration of an analyte, such as glucose, but also can be any chemical or physical property of the sample. The method of Robinson et al. involves a two-step process that includes both calibration and prediction steps. In the calibration step, the infrared light is coupled to calibration samples of known characteristic values so that there is differential attenuation of at least several wavelengths of the infrared radiation as a function of the various components and analytes comprising the sample with known characteristic value. The infrared light is coupled to the sample by passing the light through the sample or by reflecting the light from the sample.
Absorption of the infrared light by the sample causes intensity variations of the detected (passed or backscattered) light that are a function of the wavelength. The resulting intensity variations at the at least several wavelengths are measured for the set of calibration samples of known characteristic values. Original or transformed intensity variations are then empirically related to the known characteristics of the calibration samples using a multivariate algorithm to obtain a multivariate calibration model. In the prediction step, the infrared light is coupled to a sample of unknown characteristic value, and the calibration model is applied to the original or transformed intensity variations of the appropriate wavelengths of light measured from this unknown sample. The result of the prediction step is the estimated value of the characteristic of the unknown sample.
Several of the embodiments disclosed by Robinson et al. are non-invasive and incorporate an optical interface having a sensor element. As depicted in FIGS. 5 and 6 of Robinson et al., the optical interface includes an input element and an output element. The input element is an infrared light source or near infrared light source. The input element interface with the sample or body portion containing blood to be tested includes transmitting the light energy or propagating the light energy to the surface of the skin via the air. The output element includes a detector which receives the transmitted or reflected light energy. The output interface with the sample also includes propagating the transmitted or reflected light through the air from the skin.
Robinson (U.S. Pat. No. 5,830,132) discloses a robust, accurate, non-invasive analyte monitor, and is also incorporated herein by reference. The method includes irradiating the tissue with infrared energy having at least several wavelengths in a given range of wavelengths so that there is differential absorption of at least some of the wavelengths by the tissue as a function of the wavelengths and the known characteristic, wherein the differential absorption causes intensity variations of the wavelengths incident from the tissue. The method further includes providing a first path through the tissue and a second path through the tissue, wherein the first path is optimized for a first sub-region of the range of wavelengths to maximize the differential absorption by at least some of the wavelengths in the first sub-region and then optimizing the second path for a second sub-region of the range to maximize the differential absorption by at least some of the wavelengths in the second sub-region. Robinson further discloses that the object of the invention is to measure blood analytes; therefore, maximizing the amount of blood in the tissue being irradiated is recognized as improving the measurement.
The accuracy of non-invasive measurement has historically been determined by its correlation to standard invasive blood measurements. To improve the stability and accuracy of the Robinson measurement, it is disclosed that sampling device should be thermostated so that the device does not act as a heat sink. It is further disclosed that the sampling device can be heated to an above normal tissue temperature to increase blood flow to the tissue area in contact with the device. The result is an increase in the vascular supply to the tissue and a corresponding increase in the blood content of the tissue. The end result of temperature regulation is taught as a reduction in spectral variation not associated with glucose and an improvement in measurement accuracy.
An additional problem exists with respect to measuring changing blood glucose levels. While a blood draw from within a blood vessel admittedly provides an intrusive, but accurate, measure of glucose concentration, a single blood draw might not indicate whether the glucose level, at the time of the draw, is constant, rising, or dropping. A second draw, occurring shortly after the first draw, can be used to determine whether the glucose concentration is changing, as well as the rate and direction of change. However, just like the first blood draw, the second blood draw would be another discomforting inconvenience; the additional discomfort and inconvenience can further discourage patient self-monitoring. Because, as explained above, information about the direction and rate of change of glucose concentration can be critical or at least useful information, it would be advantageous to provide a system for determining, non-invasively, the direction and rate of change of analyte concentrations in tissue and/or blood. Further, it would be useful to be able to use the measured analyte concentration and rate of change to determine the analyte's concentration in other, kinetically related, compartments.
Similar problems exist with regards to measurement of other analytes. As another example, the presence or concentration of alcohol can be determined noninvasively. As with glucose, the direction and rate of change of alcohol concentration can be important in many applications. For example, the rate of change can be useful in law enforcement situations where an officer administers an alcohol measurement to a person suspected of driving under the influence. In situations where the measured alcohol concentration is near, but below, the legal limit the officer may elect to allow the driver to continue to operate their vehicle. In this scenario, the person's alcohol could be increasing such that his or her blood alcohol will be above the legal limit in a matter of minutes. Current measurement methods cannot detect these circumstances. The alcohol rate of change measurement can alter the officer's decision by conveying that the persons alcohol concentration is near the legal limit and increasing. The officer could then detain the driver and perform a second measurement after a waiting period.
Accordingly, there is a need for an apparatus and method to determine whether analyte concentrations are rising, falling or at equilibrium along with an indication of the rate of change in order to optimize treatment. The present invention addresses these needs as well as other problems associated with existing methods for non-invasively measuring levels of and changes in analyte concentration in blood utilizing spectroscopy, and further applies to non-invasive measurements of analyte concentrations for other analytes and solutions.